Ventilatory assist devices are machines used in the treatment of respiratory failure and sleep disorders in hospital or home settings. With assisted ventilation (e.g. assist volume cycled ventilation, pressure support ventilation, bi-level assist in the case of non-invasive devices and proportional assist ventilation) ventilator cycles are triggered by the patient and are intended to coincide with patient's inspiratory effort, beginning the support when inspiratory effort starts and ending the support at the end of patient's inspiratory effort. In practice, however, the ventilator cycle never begins at the onset of patient's inspiratory effort (trigger delay) and the end of the ventilator's inflation phase only rarely coincides with the end of inspiratory effort (cycling-off errors). FIG. 1 provides an example. The bottom channel is transdiaphragmatic pressure (measured by esophageal and gastric catheters) and reflects true patient inspiratory effort. As may be seen, ventilator cycle was triggered several hundred milliseconds after onset of effort (interval between vertical lines) and the inflation cycle continued well beyond the effort. In fact, the ventilator was cycling almost completely out-of-phase with the patient. Trigger delay is often so marked that some efforts completely fail to trigger the ventilator (ineffective efforts, e.g. third effort, FIG. 1). A more advanced form of non-synchrony is shown in FIG. 2. In this case, the inflation cycle of the ventilator extends over two patient cycles. There are, accordingly, two inspiratory efforts within a single inflation phase and there is an additional ineffective effort during the ventilator's expiratory phase. The arrows in FIG. 2 indicate the location of the extra patient efforts that did not trigger corresponding ventilator cycles.
Non-synchrony between patient and ventilator is extremely common. Leung et al found that, on average, 28% of patient's efforts are ineffective (Leung P, Jubran A, Tobin M J (1997). Comparison of assisted ventilator modes on triggering, patient effort, and dyspnea. Am J Respir Crit Care Med 155:1940-1948). Considering that ineffective efforts are the extreme manifestation of non-synchrony, less severe, yet substantial (e.g. first two breaths, FIG. 1), delays must occur even more frequently. Non-synchrony is believed to cause distress, leading to excessive sedation and sleep disruption, as well as errors in clinical assessment of patients since the respiratory rate of the ventilator can be quite different from that of the patient. Monitoring respiratory rate is a fundamental tool for monitoring critically ill patients on ventilators. Non-synchrony is not only prevalent in intensive care units but is also frequently present in the home setting during sleep when patients are receiving bi-level support for the treatment of sleep apnea or respiratory failure (personal observations). The present invention concerns a novel method and apparatus to, non-invasively, automatically and in real-time, generate a signal that reflects changes in inspiratory effort. Such a signal can then be used, among other things, to determine the true onset (Tonset) and end (Tend) of patient's inspiratory efforts. Such method/device can be used simply as a monitor, informing the user of the presence, manifestations and magnitude of non-synchrony. The user can then take appropriate action to reduce the non-synchrony. Alternatively, the method/device can be coupled with the ventilator's cycling mechanisms, whereby onset and end of ventilator cycles are automatically linked to onset and end of patient's efforts, thereby insuring synchrony without intervention by the user.
In current ventilatory assist devices, triggering usually occurs when flow becomes inspiratory (i.e. >0) and exceeds a specified amount, or when airway pressure decreases below the set PEEP (positive end-expiratory pressure) level by a specified amount. Trigger delay has two components. One component is related to ventilator trigger response and sensitivity. Thus, if the response of the ventilator is poor, triggering may not occur immediately when the triggering criteria are reached. Alternatively, the threshold for triggering may be set too high by the user. The component of trigger delay attributable to ventilator response and sensitivity is given by the interval between zero flow crossing (arrow, FIG. 1) and triggering (second vertical line). The response of modern ventilators has improved substantially over the past several years such that it is difficult to effect further improvements in this respect, and this invention does not contemplate any such improvements. This component of trigger delay can, however, still be excessive if the user sets an unnecessarily high threshold. This setting may be because of lack of sufficient expertise, or because there was excessive baseline noise at some point, which necessitated a high threshold to avoid auto-triggering. The threshold then remains high even after disappearance of the noise.
The second component of trigger delay is the time required, beyond the onset of inspiratory effort (Tonset), for expiratory flow to be reduced to zero (interval between first vertical line and the arrow, FIG. 1). This delay is related to the fact that expiratory resistance is usually high in ventilated patients and expiratory time is frequently too short to allow lung volume to return to FRC (functional residual capacity) before the next effort begins. At Tonset, therefore, elastic recoil pressure is not zero (DH, dynamic hyperinflation). Inspiratory effort must first increase enough to offset the elastic recoil pressure associated with DH before flow can become inspiratory, and/or before Paw (airway pressure) decreases below PEEP, in order to trigger the ventilator. By identifying the true Tonset, a capacity that is permitted by current invention, this component of trigger delay (usually the largest component, seen, for example, FIG. 1) can be essentially eliminated.
Cycling-off errors result from the fact that, except with Proportional Assist Ventilation, current ventilator modes do not include any provision that links the end of ventilator cycle to end of the inspiratory effort of the patient. In the most common form of assisted ventilation, Volume Cycled Ventilation, the user sets the duration of the inflation cycle without knowledge of the duration of patient's inspiratory effort. Thus, any agreement between the ends of ventilator and patient inspiratory phases is coincidental. With the second most common form, Pressure Support Ventilation, the inflation phase ends when inspiratory flow decreases below a specified value. Although the time at which this threshold is reached is, to some extent, related to patient effort, it is to the largest extent related to the values of passive resistance and elastance of the patient. In patients in whom the product [resistance/elastance], otherwise known as respiratory time constant, is high, the ventilator cycle may extend well beyond patient effort, while in those with a low time constant the cycle may end before the end of patient's effort (Younes M (1993) Patient-ventilator interaction with pressure-assisted modalities of ventilatory support. Seminars in Respiratory Medicine 14:299-322; Yamada Y, Du H L (2000) Analysis of the mechanisms of expiratory asynchrony in pressure support ventilation: a mathematical approach. J Appl Physiol 88:2143-2150). By providing a signal that reflects changes in inspiratory effort, the current invention makes it possible to determine when effort begins declining, thereby making it possible to synchronize the end of ventilator cycle with end of patient's effort.
In U.S. Pat. No. 6,305,374 B1, an approach is described to identify the onset and end of patient's inspiratory effort during non-invasive bi-level positive pressure ventilation (BiPAP). This approach relies exclusively on the pattern of flow waveform to make these identifications. Thus, current values of flow are compared with an estimated value based on projections from preceding flow pattern. If the difference exceeds a preset amount, a phase switch is declared. There is no attempt whatsoever in this method to generate a signal that continuously reflects the pattern of inspiratory effort in real-time throughout the breath. Furthermore, while this method may yield reasonably accurate results in the intended application (treatment of obstructive sleep apnea patients with non-invasive BiPAP), a number of considerations suggest that its use in critically ill, intubated, ventilated patients may not provide accurate results:
1) Implicit to the use of flow as a marker of respiratory muscle pressure output is the assumption that flow pattern reflects changes in alveolar pressure inside patient's lung. This is where respiratory muscle pressure is exerted. This assumption, however, is true only if airway pressure is constant. Since airway pressure is one of the two pressure values that determine flow (flow=(airway pressure-alveolar pressure)/resistance), it is clear that changes in airway pressure can alter flow even if there is no change in respiratory muscle or alveolar pressure. In non-invasive bi-level support, airway pressure, one of the two pressure values that determine flow, is reasonably constant during both inspiration and expiration, even though the absolute level is different in the two phases. If one of the two pressure values is constant during a given phase, it is reasonable to assume that changes in flow during that phase reflect changes in the other pressure, namely alveolar pressure. This condition does not apply in intubated, mechanically ventilated patients. In most modern intensive care ventilators, airway pressure is actively controlled during expiration through adjustments of the PEEP/exhalation valve mechanism. The pattern of such active changes in airway pressure during expiration varies from one ventilator brand to another and in the same ventilator from time to time depending on the state of the PEEP/exhalation valve mechanism. Under these conditions, changes in flow trajectory during expiration cannot be assumed to reflect changes in alveolar pressure trajectory. Likewise, during inspiration airway pressure is far from being constant, regardless of the mode used. Thus, changes in inspiratory flow profile cannot be used to reflect similar changes in alveolar pressure. The use of flow to infer end of effort during the inflation phase is accordingly not plausible.
2) When passive elastance (E) and resistance (R) are constant over the entire tidal volume range, the product R/E, or respiratory time constant, is also constant over the entire period of expiration. Because the time constant governs the pattern of lung emptying, a constant R/E produces a predictable exponential flow pattern in the passive system. With a predictable pattern it is possible to make forward extrapolations, or predictions, for the sake of identifying a deviation from the expected passive behaviour. Such deviation may then be used, with reasonable confidence, to infer the development of an additional active force, such as the onset of inspiratory muscle effort. When E and R are not constant throughout the breath, R/E may change from time to time causing changes in flow trajectory (Δflow/Δt) that are not related to muscle pressure. Under these conditions, deviation in Δflow/Δt from previous values cannot reliably signify a change in pressure generated by respiratory muscles. Patients with obstructive sleep apnea, the intended population of U.S. Pat. No. 6,305,374 B1, have generally normal lungs; R and E are expected to be constant over the tidal volume range, particularly when expiratory airway pressure is higher than atmospheric (i.e. the usual case when BiPAP is applied). In critically ill, intubated ventilated patients, this is not the case. Resistance is not constant, primarily because these patients are intubated and the resistance of the endotracheal tube is flow-dependent (the higher the flow, the higher the resistance). The relation between resistance and flow varies from one tube to the other. Furthermore, tidal volume in these patients often extends into the volume range where elastance is not constant. Thus, as the lung is emptying, either or both elastance and resistance may be changing, causing changes in respiratory time constant during the same expiration. Under these conditions, changes in flow trajectory need not reflect changes in respiratory muscle pressure. This considerably decreases the sensitivity and specificity of flow pattern as a marker of inspiratory effort.
3) Changes in respiratory muscle pressure (Pmus) are not exclusively used to change flow. According to the equation of motion, specifically applied to intubated patients:Pmus=Volume*E+Flow*K1+(Flow*absolute flow*K2)−Paw  Equation 1
Where, E is passive respiratory system elastance, K1 is the laminar component of passive respiratory system resistance, K2 is the resistance component related to turbulence (mostly in the endotracheal tube or nasal passages), and Paw is airway pressure which is determined by the pressure at the exhalation/PEEP valve (Pvalve), flow and Rex, that is resistance of the exhalation tubing (Paw=Pvalve−flow*Rex). In this equation expiratory flow is negative. When Pmus changes, as at Tonset, the flow trajectory should change. However, a change in flow trajectory also results in changes in volume and Paw trajectories. According to Equation 1, these changes will oppose the change in flow. For example, if expiratory flow decreases at a faster rate, volume decreases at a slower rate than in the absence of Pmus. At any instant after Tonset, elastic recoil pressure, which is related to volume, is higher, and this promotes a greater expiratory flow. The same can be said for the effect of changes in flow trajectory on Paw trajectory; a lower expiratory flow decreases Paw, which promotes more expiratory flow. How much of the change in Pmus is used to change the flow trajectory depends on the magnitude of the opposing forces. In particular, a higher passive elastance and/or a higher Rex tends to reduce the fraction of the change in Pmus used to change flow trajectory. Furthermore, for a given Pmus expended to change the flow trajectory, the actual change in trajectory is determined by resistance (i.e. K1 and K2). When E, Rex, K1 and K2 are all low, a modest change in dPmus/dt results in a sharp change in flow trajectory. As these characteristics become more abnormal, the change in flow trajectory, for a given dPmus/dt, progressively is attenuated. FIG. 3 illustrates this in a computer simulation.
In the example of FIG. 3, respiratory muscles were inactive in the first second of expiration (as they usually are). This is represented by Pmus of zero (lower panel). At 1.0 sec an inspiratory effort begins. Pmus rises at a rate of 10 cmH2O/sec, representative of a normal respiratory drive. The three flow waveforms represent, from below upwards, progressively increasing values of K1, K2, E and Rex. The values used in the lowest waveform are those of a patient with normal passive elastance and resistance, intubated with a large endotracheal tube (#9 tube, K2=3), and exhalation tubing with a low resistance (Rex=2). The onset of effort results in a sharp change in the flow trajectory that can be readily detected within a very short time after Tonset.
The middle waveform (FIG. 3) was generated with values representing the average intensive care patient on mechanical ventilation. Both passive K1 and passive E are higher than normal, K2 is that of a #8 endotracheal tube, the most common size used, and the exhalation tubing has a moderate (average) resistance. Note that the change in flow trajectory is considerably less pronounced. An experienced eye, with the benefit of hindsight (i.e. observing the flow waveform for a substantial period after Pmus started), may be able to tell that a change in trajectory occurred at 1.0 sec. However, it is not possible to prospectively identify that a trajectory change took place in a timely manner, for the sake of triggering the ventilator. Prospective identification of a trajectory change requires comparison between current and previous Δflow/Δt values, or between current flow values and values expected based on forward extrapolation of the preceding flow pattern (e.g. dashed lines, FIG. 3). There is always uncertainty with extrapolation, particularly with non-linear functions where the exact function is not known and, even more so, when the signal is noisy, as the flow signal commonly is (due to cardiac artefacts or secretions). Comparison of current and previous Δflow/Δt is also fraught with uncertainties when the rate may change for reasons other than respiratory muscle action (see #1 and #2, above). Thus, a wide difference (trigger threshold) must be specified, between current and projected flow, or between current and previous Δflow/Δt, before a trajectory change can be identified with confidence. Otherwise, false triggering will occur frequently. When the change in flow trajectory is small, a longer interval must elapse before the threshold separation is achieved. It can be seen from the middle flow waveform that a conservative flow separation (between actual and projected flow) of 0.2 l/sec would not be reached until after flow became inspiratory. Thus, in the average mechanically ventilated patient the use of flow trajectory to identify Tonset is not likely to result in a significant improvement over the current approach of waiting for flow to become inspiratory.
With more severe mechanical abnormalities (top waveform, FIG. 3), the change in flow trajectory is even more subtle. Even an experienced eye, with the benefit of hindsight, cannot distinguish between a true trajectory change and some flow artefact. Clearly, with a much stronger effort a flow trajectory change may be identifiable before flow becomes inspiratory. However, when patients have vigorous inspiratory efforts, there is no significant trigger delay even with current triggering techniques.
In summary, the use of flow to identify respiratory phase transitions is entirely unsuitable for identification of inspiratory to expiratory transitions during mechanical ventilation in critically ill patients (because of the highly variable Paw during inflation), and has poor sensitivity and specificity for identifying expiratory to inspiratory transitions in these patients because of the frequent use of active exhalation valves, the presence of variable time constant during expiration and the often marked abnormalities in elastance and resistance.
An alternative approach has recently been proposed by Younes (U.S. patent application Ser. No. 10/517,384 filed Dec. 10, 2004, the disclosure of which is incorporated herein by reference and corresponding EP application 03 739906 filed Jun. 27, 2003 (WO 2003/002561); Method and Device for monitoring and Improving Patient-Ventilator Interaction). The approach consists of generating a PMUS waveform using improvised values of elastance and resistance. Here, the above equation 1 is used to generate PMUS but, instead of using real resistance (K1) and elastance (E) values, which are difficult to obtain in spontaneously breathing patients, improvised values are used which simply result in the generated PMUS waveform having the shape characteristics of normally occurring PMUS waveforms, namely an approximately flat baseline during expiration and a ramp-like rising phase in the inspiratory phase. The surrogate values for elastance and resistance are assigned herein, the terms KV and KF to distinguish them from the real values. Once such an improvised PMUS signal is generated, it is possible to easily identify the onsets and ends of inspiratory efforts for the sake of triggering and cycling-off ventilators. Because the PMUS generated by these improvised resistance and elastance values is not a real PMUS signal, the value generated by the current approach is referred herein to simply as Signal.
The above invention described by Younes proposes the use of a default value for KF and adjusting the KV value to result in a flat baseline during expiration. Alternatively, a default value for KV is used while the KF value is adjusted to result in a flat baseline during expiration. The preferred embodiments in this earlier Younes patent application employ a fixed value for one of the two variables while adjusting the value of the other variable manually with visual feedback from a monitor. Although the specification suggests that appropriate values for KV and KF may be selected automatically using appropriate software, the specification does not teach any approach for doing that and it is evident that such software would have to be sufficiently sophisticated to replace the complex functions executed by the eye-brain combination in humans.
The present invention proposes new methods and apparatus to supplement the approach proposed by Younes. These improvements relate to methods for automatically (as opposed to manually) determining the values of KF and KV required for generating a physiologically appropriate Signal waveform from which information about onsets and ends of inspiratory efforts can be derived. Specifically, these methods employ complex algorithms to distinguish between true baseline and noise values during expiration, a task that can be readily done by the human eye, but is very difficult to translate into computer instructions.
Because these new methods/device are intended to work with, and represent an improvement over, the original Younes approach the latter approach will be described in some detail in the detailed description of the invention, below.